Data acquisition method and apparatus for MR imaging

ABSTRACT

An improved technique is described for acquiring MR image data such as needed for FSE-based Dixon imaging techniques. A gradient-induced echo shift is produced in the pulse sequences by a small gradient applied along the readout axis prior to a readout pulse. When necessary, another small pulse is applied along the readout axis, equal in area and opposite in polarity to the first, to compensate for the shifting effect. Similar pulses are applied for each acquisition window. While data with non-zero phase shifts between water and fat signals are collected as fractional echoes, no increase in echo spacing is necessary with the modified acquisition strategy. Images corresponding to different phase shifts are reconstructed using phase-sensitive partial Fourier reconstruction algorithms whenever necessary. These images are then used to separate different chemical species (such as water and fat) in the object to be imaged. Increased time efficiency is therefore achieved with the improved technique, with a significant reduction in degradation due to losses in slice coverage and increased image blurring and sensitivity to flow and motion due to T2-modulation along the echo train in conventional techniques.

BACKGROUND OF THE INVENTION

[0001] The present invention relates generally to the field of magneticresonance imaging. More particularly, the invention relates to atechnique for providing magnetic resonance echo shifting without theneed for shifting the radio-frequency pulse, or the data acquisitionwindow, or both, such as for producing water-only or fat-only images, ina highly time efficient manner through a Dixon imaging pulse sequence.

[0002] Magnetic resonance imaging (MRI) systems have become ubiquitousin the field of medical diagnostics. Over the two past decades, improvedtechniques for MRI examinations have been developed that now permit veryhigh-quality images to be produced in a relatively short time. As aresult, diagnostic images with varying degrees of resolution areavailable to the radiologist that can be adapted to particulardiagnostic applications.

[0003] In general, MRI examinations are based on the interactions amonga primary magnetic field, a radiofrequency (rf) magnetic field and timevarying magnetic gradient fields with nuclear spins within the subjectof interest. Specific nuclear components, such as hydrogen nuclei inwater molecules, have characteristic behaviors in response to externalmagnetic fields. The precession of spins of such nuclear components canbe influenced by manipulation of the fields to produce rf signals thatcan be detected, processed, and used to reconstruct a useful image.

[0004] The magnetic fields used to generate images in MRI systemsinclude a highly uniform, static magnetic field that is produced by aprimary magnet. A series of gradient fields are produced by a set ofthree gradient coils disposed around the subject. The gradient fieldsencode positions of individual volume elements or voxels in threedimensions. An rf coil is employed to produce an rf magnetic field. Thisrf magnetic field perturbs the spin system from its equilibriumdirection, causing the spins to precess around the axis of theirequilibrium magnetization. During this precession, rf fields are emittedby the spins and are detected by either the same transmitting rf coil,or by a separate receive-only coil. These signals are amplified,filtered, and digitized. The digitized signals are then processed usingone of several possible reconstruction algorithms to reconstruct auseful image.

[0005] Many specific techniques have been developed to acquire MR imagesfor a variety of applications. One major difference among thesetechniques is in the way gradient pulses and rf pulses are used tomanipulate the spin systems to yield different image contrasts,signal-to-noise ratios, and resolutions. Graphically, such techniquesare illustrated as “pulse sequences” in which the pulses arerepresented, along with temporal relationships among them. In recentyears, pulse sequences have been developed which permit extremely rapidacquisition of large amounts of raw data. Such pulse sequences permitsignificant reduction in the time required to perform the examinations.Time reductions are particularly important for acquiring high resolutionimages, as well as for suppressing motion effects and reducing thediscomfort of patients in the examination process.

[0006] Among the pulse sequences which have been developed for fastacquisition of large amounts of MR data, is a sequence generallyreferred to as fast spin echo (FSE). This technique is capable ofgenerating high-quality image data in a fraction of the time needed forconventional spin echo imaging. FSE techniques have thus become thesequence of choice, especially for T2-weighted imaging. However, aprominently distinguishing feature of FSE images is an anomalouslybright signal resulting from fat content in the tissue being imaged. Thephenomenon has been attributed to the demodulation of the J-coupling andde-sensitization of diffusion through inhomogeneities due to the rapidlyrefocusing rf pulse trains contained in the FSE pulse sequence.

[0007] Fat suppression has therefore become desirable in T2-weightedimaging procedures. At present, several techniques have been employedfor such fat suppression. A first such technique is referred to aschemical saturation, and can be used to reduce the fat signal, butrequires very homogeneous magnetic fields due to the close separation ofthe water and fat signals resulting from the excitation. In particular,the rf pulse must saturate all fat, requiring a highly uniform mainmagnetic field, to avoid separating water signals. Similarly, thetechnique depends highly upon the homogeneity of the rf field which isneeded to achieve an accurate flip of the fat signal for suppression andsubsequent flip of the resulting water signals for imaging.Inhomogeneity in the main magnetic field is particularly a problem atlocations off the isocenter of the field system. Finally, patientanatomy also tends to perturb the fields, rendering the techniqueparticularly problematic.

[0008] A second technique that has been developed for fat suppressioninvolves short inversion time (TI) inversion recovery, and is commonlyreferred to as STIR. This technique is intended to flip all signals toan inverted direction, with fat and water signals recovering atdifferent rates. The technique then acquires the image data when the fatsignal is crossing the null point while the water signal is stillpartially in the inverted state. Because of its underlying principles,the technique typically is dependent on the Ti of the water signal, andgenerally results in relatively low signal-to-noise ratios due to thepartial recovery of the water signal during the recovery of the fatsignal.

[0009] A further technique that has been developed is generally referredto as the Dixon technique. In this approach, the chemical shiftdifference between water and fat is encoded into images with differentecho shifts. Field inhomogeneity effects appear as image phase errors,which in principle can be corrected for by a combination of multipointacquisition and more elaborate image processing. While these techniquesallow for more uniform water and fat separation in the presence of fieldinhomogeneity, one clear drawback is the requirement for multiple dataacquisitions and therefore longer scan times.

[0010] Incorporating the Dixon approach with fat suppression into FSEpulse sequences presents a mutually beneficial combination. While theDixon technique provides a potentially robust separation of the strongfat signal, FSE helps to alleviate for long data acquisition times inthe multipoint Dixon technique. In an exemplary combination of thesetechniques, however, echo shift as dictated by the Dixon technique wasachieved by shifting the timing of the readout gradient and the dataacquisition window to maintain necessary conditions(Carr-Purcell-Meiboom-Gill; “CPMG” conditions). As a result, inter-echospacing was increased, leading to substantial loss in the slice coveragefor a given sequence repetition time, largely offsetting the gain ofusing FSE for reducing the scan time. The technique is believed,therefore, to be appropriate for imaging small anatomic areas only thatdo not require large slice coverage.

[0011] Dixon technique based on the conventional spin echo or gradientecho sequences generally employ shifting the echo through eithershifting the RF pulse, or the data acquisition window, or both. As inthe case of FSE based Dixon technique, such shifting may lead to thedisadvantage of longer acquisition times because of the increaseddeadtime during a sequence. Consequently, the loss of slice coverage fora given span time, or increased scan time for a given number of slices,and an increase in blurring and greater sensitivity to flow and motionartifacts, can all result.

[0012] There is a need, therefore, for an improved technique forobtaining shifts in echos in MR imaging sequences. There is a particularneed for a FSE-based Dixon imaging approach which achieves the echoshifts satisfying the CPMG conditions without necessitating an increasein echo spacing. There is, at present, a particular need for an improvedtechnique which can be implemented on existing hardware and controlsystems to obtain the improvement in timing and imaging clarity in arelatively straightforward manner.

BRIEF DESCRIPTION OF THE INVENTION

[0013] The present invention provides an imaging technique designed torespond to such needs. In accordance with one aspect of the technique, amethod is provided for acquiring magnetic resonance image data. Themethod includes a step of, in the presence of a primary and gradientmagnetic field system, generating an echo shifting gradient pulse on areadout access. A readout gradient pulse is then generated on thereadout axis. Magnetic resonance echo signals are detected that resultfrom the readout gradient, and a compensating gradient pulse is thengenerated on the readout axis.

[0014] In accordance with another aspect of the invention, a method issimilarly provided for acquiring magnetic resonance image data thatincludes, in the presence of a primary and gradient field magnetic fieldsystem, applying a Dixon fast spin echo pulse sequence. An echo shiftinggradient pulse is generated on a readout axis, and a readout gradientpulse is then generated on the readout axis. Magnetic resonance echosignals resulting from the readout gradient are detected, and acompensating gradient pulse is generated on the readout axis.

[0015] In accordance with another aspect of the invention, a method foracquiring magnetic resonance image data includes apply a Dixon fast spinecho pulse sequence to acquire a plurality of k-space lines of data inthe presence of a primary and gradient magnetic field system. For eachk-space line of data, an echo shifting gradient pulse of a firstpolarity and of a desired area is generated on a readout axis. A readoutgradient pulse is generated on the readout axis as well, and magneticresonance echo signals resulting from the readout gradient are detected.Finally, compensating gradient pulses are generated of a second polarityopposite to the first polarity and of the desired area on the readoutaxis.

[0016] In accordance with the further aspect of the invention, a methodfor acquiring magnetic resonance image data includes applying a Dixonfast spin echo pulse sequence to acquire a plurality of k-space lines ofdata in the presence of a primary and gradient magnetic field system.For each k-space line of data, an echo shifting gradient pulse of afirst polarity and of a desired area is generated on a readout axis, anda readout gradient pulse is generated on the readout axis. Magneticresonance echo signals resulting from the readout gradient are detected,and a compensating gradient pulse of a second polarity opposite to thefirst polarity, and of the desired area is generated on a readout axis.An image is then reconstructed based upon the detected echo signals.

BRIEF DESCRIPTION OF THE DRAWINGS

[0017]FIG. 1 is a diagrammatical representation of an MRI system for usein medical diagnostic imaging and implementing certain aspects of thepresent shielding technique;

[0018]FIG. 2 is diagrammatical representation of a pulse sequencedescription for a conventional Dixon fast spin echo MRI sequence,particularly illustrating time delays required for the Dixon technique;

[0019]FIG. 3 is a diagrammatical representation of a pulse sequencedescription in accordance with aspects of the present technique forshifting echo times to reduce the sequential and overall time requiredfor signal acquisition in an MRI sequence of the Dixon fast spin echotype;

[0020]FIG. 4 is a diagrammatical representation of a pulse sequencedescription similar to that of FIG. 3 but wherein echo shifting gradientpulses have been inverted in polarity; and

[0021]FIG. 5 is diagrammatical representation of an alternative pulsesequence description in accordance with aspects of the presenttechnique.

DETAILED DESCRIPTION OF SPECIFIC EMBODIMENTS

[0022] Turning now to the drawings, and referring first to FIG. 1, amagnetic resonance imaging (MRI) system 10 is illustrateddiagrammatically as including a scanner 12, scanner control circuitry14, and system control circuitry 16. While MRI system 10 may include anysuitable MRI scanner or detector, in the illustrated embodiment thesystem includes a full body scanner comprising a patient bore 18 intowhich a table 20 may be positioned to place a patient 22 in a desiredposition for scanning. Scanner 12 may be of any suitable type of rating,including scanners varying from 0.5 Tesla ratings to 1.5 Tesla ratingsand beyond.

[0023] Scanner 12 includes a series of associated coils for producingcontrolled magnetic fields, for generating radiofrequency rf excitationpulses, and for detecting emissions from gyromagnetic material withinthe patient in response to such pulses. In the diagrammatical view ofFIG. 1, a primary magnet coil 24 is provided for generating a primarymagnetic field generally aligned with patient bore 18. A series ofgradient coils 26, 28 and 30 are grouped in a coil assembly forgenerating controlled magnetic gradient fields during examinationsequences as described more fully below. An rf coil 32 is provided forgenerating rf pulses for exciting the gyromagnetic material. Power issupplied to scanner 12 in any appropriate manner, as indicated generallyat reference numeral 34. In the embodiment illustrated in FIG. 1, coil32 also serves as a receiving coil. Thus, rf coil 32 may be coupled withdriving and receiving circuitry in passive and active modes forreceiving emissions from the gyromagnetic material and for applying rfexcitation pulses, respectively. Alternatively, various configurationsof receiving coils may be provided separate from rf coil 32. Such coilsmay include structures specifically adapted for target anatomies, suchas head coil assemblies, and so forth. Moreover, receiving coils may beprovided in any suitable physical configuration, including phased arraycoils, and so forth. As described more fully below, the presenttechnique provides for improved image acquisition sequences whichenhance image quality, particularly in the separation of signals fromdiffering materials, such as water and fat, without significantlengthening in image data acquisition times.

[0024] In a present configuration, the gradient coils 26, 28 and 30 havedifferent physical configurations adapted to their function in theimaging system 10. As will be appreciated by those skilled in the art,the coils are comprised of conductive wires, bars or plates which arewound or cut to form a coil structure which generates a gradient fieldupon application of control pulses as described below. The placement ofthe coils within the gradient coil assembly may be done in severaldifferent orders, but in the present embodiment, a Z-axis coil ispositioned at an innermost location, and is formed generally as asolenoid-like structure which has relatively little impact on the rfmagnetic field. Thus, in the illustrated embodiment, gradient coil 30 isthe Z-axis solenoid coil, while coils 26 and 28 are Y-axis and X-axiscoils respectively.

[0025] The coils of scanner 12 are controlled by external circuitry togenerate desired fields and pulses, and to read signals from thegyromagnetic material in a controlled manner. As will be appreciated bythose skilled in the art, when the material, typically bound in tissuesof the patient, is subjected to the primary field, individual magneticmoments of the paramagnetic nuclei in the tissue partially align withthe field. While a net magnetic moment is produced in the direction ofthe polarizing field, the randomly oriented components of the moment ina perpendicular plane generally cancel one another. During anexamination sequence, an rf frequency pulse is generated at or near theLarmor frequency of the material of interest, resulting in rotation ofthe net aligned moment to produce a net transverse magnetic moment. Thistransverse magnetic moment precesses around the main magnetic fielddirection, emitting rf signals that are detected by the scanner andprocessed for reconstruction of the desired image.

[0026] Gradient coils 26, 28 and 30 serve to generate preciselycontrolled magnetic fields, the strength of which vary over a predefinedfield of view, typically with positive and negative polarity. When eachcoil is energized with known electric current, the resulting magneticfield gradient is superimposed over the primary field and produces adesirably linear variation in the Z-axis component of the magnetic fieldstrength across the field of view. The field varies linearly in onedirection, but is homogenous in the other two. The three coils havemutually orthogonal axes for the direction of their variation, enablinga linear field gradient to be imposed in an arbitrary direction with anappropriate combination of the three gradient coils.

[0027] The pulsed gradient fields perform various functions integral tothe imaging process. Some of these functions are slice selection,frequency encoding and phase encoding. These functions can be appliedalong the X-, Y- and Z-axis of the original coordinate system or alongother axes determined by combinations of pulsed currents applied to theindividual field coils.

[0028] The slice select gradient determines a slab of tissue or anatomyto be imaged in the patient. The slice select gradient field may beapplied simultaneously with a frequency selective rf pulse to excite aknown volume of spins within a desired slice that precess at the samefrequency. The slice thickness is determined by the bandwidth of the rfpulse and the gradient strength across the field of view.

[0029] The frequency encoding gradient is also known as the readoutgradient, and is usually applied in a direction perpendicular to theslice select gradient. In general, the frequency encoding gradient isapplied before and during the formation of the MR echo signal resultingfrom the rf excitation. Spins of the gyromagnetic material under theinfluence of this gradient are frequency encoded according to theirspatial position along the gradient field. By Fourier transformation,acquired signals may be analyzed to identify their location in theselected slice by virtue of the frequency encoding.

[0030] Finally, the phase encode gradient is generally applied beforethe readout gradient and after the slice select gradient. Localizationof spins in the gyromagnetic material in the phase encode direction isaccomplished by sequentially inducing variations in phase of theprecessing protons of the material using slightly different gradientamplitudes that are sequentially applied during the data acquisitionsequence. The phase encode gradient permits phase differences to becreated among the spins of the material in accordance with theirposition in the phase encode direction.

[0031] As will be appreciated by those skilled in the art, a greatnumber of variations may be devised for pulse sequences employing theexemplary gradient pulse functions described above, as well as othergradient pulse functions not explicitly described here. Moreover,adaptations in the pulse sequences may be made to appropriately orientboth the selected slice and the frequency and phase encoding to excitethe desired material and to acquire resulting MR signals for processing.

[0032] The coils of scanner 12 are controlled by scanner controlcircuitry 14 to generate the desired magnetic field and radiofrequencypulses. In the diagrammatical view of FIG. 1, control circuitry 14 thusincludes a control circuit 36 for commanding the pulse sequencesemployed during the examinations, and for processing received signals.Control circuit 36 may include any suitable programmable logic device,such as a CPU or digital signal processor of a general purpose orapplication-specific computer. Control circuit 36 further includesmemory circuitry 38, such as volatile and non-volatile memory devicesfor storing physical and logical axis configuration parameters,examination pulse sequence descriptions, acquired image data,programming routines, and so forth, used during the examinationsequences implemented by the scanner.

[0033] Interface between the control circuit 36 and the coils of scanner12 is managed by amplification and control circuitry 40 and bytransmission and receive interface circuitry 42. Circuitry 40 includesamplifiers for each gradient field coil to supply drive current to thefield coils in response to control signals from control circuit 36.Interface circuitry 42 includes additional amplification circuitry fordriving rf coil 32. Moreover, where the rf coil serves both to emit therf excitation pulses and to receive MR signals, circuitry 42 willtypically include a switching device for toggling the rf coil betweenactive or transmitting mode, and passive or receiving mode. A powersupply, denoted generally by reference numeral 34 in FIG. 1, is providedfor energizing the primary magnet 24. Finally, circuitry 14 includesinterface components 44 for exchanging configuration and image data withsystem control circuitry 16. It should be noted that, while in thepresent description reference is made to a horizontal cylindrical boreimaging system employing a superconducting primary field magnetassembly, the present technique may be applied to various otherconfigurations, such as scanners employing vertical fields generated bysuperconducting magnets, permanent magnets, electromagnets orcombinations of these means.

[0034] System control circuitry 16 may include a wide range of devicesfor facilitating interface between an operator or radiologist andscanner 12 via scanner control circuitry 14. In the illustratedembodiment, for example, an operator controller 46 is provided in theform of a computer work station employing a general purpose orapplication-specific computer. The station also typically includesmemory circuitry for storing examination pulse sequence descriptions,examination protocols, user and patient data, image data, both raw andprocessed, and so forth. The station may further include variousinterface and peripheral drivers for receiving and exchanging data withlocal and remote devices. In the illustrated embodiment, such devicesinclude a monitor 48, a conventional computer keyboard 50 and analternative input device such as a mouse 52. A printer 54 is providedfor generating hard copy output of documents and images reconstructedfrom the acquired data. A computer monitor 48 is provided forfacilitating operator interface. In addition, system 10 may includevarious local and remote image access and examination control devices,represented generally by reference numeral 56 in FIG. 1. Such devicesmay include picture archiving and communication systems, teleradiologysystems, and the like.

[0035] Particular pulse sequence descriptions have been developed forimplementation on MRI systems of the type illustrated in FIG. 1 whichpermit acquisition of large amounts of data in relatively short timeperiods to produce high-quality reconstructed images. One such techniqueis known as fast spin echo (FSE) imaging. In the FSE technique, basedupon techniques known as relaxation enhancement (RARE) and spin echo(SE) imaging, data representative of all MR signal echoes needed forimage reconstruction are not collected in a single shot. Instead,multiple echo signal data sets are collected, typically in excess of 16,each encoded at a different phase level. The data is collected at eachrelaxation time interval, with refocusing pulses being applied by the rfcoil at repeated integrals. A major benefit in FSE and similar imagingtechniques is a great reduction in the overall time required for signalacquisition.

[0036] In order to achieve better fat suppression in the presence ofmain and/or RF field inhomogeneities, a technique commonly referred toas the Dixon method obtains echo shifts (shifts in the time ofoccurrence of the MR echo signal) by time shifting the refocusing rfpulse. Alternatively, similar shifts can be achieved by shifting thedata acquisition window and the readout gradient while keeping thetiming of the refocusing pulse fixed. The amount of the time shift ineither cases is directly proportional to the desired phase shift betweendifferent paramagnetic species intended to be separated for imagereconstruction. As noted above, it is particularly useful to separatewater and fat magnetizations in imaging sequences so as to more clearlyrepresent anatomical features of interest. FIG. 2 represents anexemplary FSE implementation of the Dixon technique using the latterapproach of shifting the readout gradient and the data acquisitionwindow of this type.

[0037] As shown in FIG. 2, the pulse sequence description, referred togenerally by reference numeral 58, may be represented diagrammaticallyas a series of pulses applied on logical axis of the MRI system. As willbe appreciated by those skilled in the art, the logical axis correspondsto activities imposed on the various system components, particularly thegradient and rf coils. In practice, the pulses indicated the presentdiscussion as being applied to a readout axis may, of course, be appliedto one or multiple physical axes defined by the gradient coils.Similarly, other conventional pulses will be included in the pulsesequences and applied to the gradient coils for phase and frequencyencoding.

[0038] In the diagrammatical representation of FIG. 2, the pulsesequence 58 is illustrated on three logical axes, the rf axis 60, thereadout gradient axis 62, and the MR signal or data acquisition axis 64.In the illustration of FIG. 2, a first 90° excitation pulse 66 isapplied to the rf axis, followed by a dephaser pulse 68 applied on thereadout axis 62. Following a predetermined time period, a 1800 refocuspulse 70 is applied to the rf axis 60. Thereafter, a readout pulse 72 isapplied along the readout axis 62, to collect signal data for an MR echo78 occurring as indicated along data acquisition axis 64. The pulsesresult in different-echo times as indicated by lines 74 and 76,respectively, in FIG. 2. These echo times correspond to echoes withdifferent phase shifts from the different constituents for which imagedata is desired, such as water and fat. A time delay 80 is therebyintroduced by the shift which recurs for each repetition of dataacquisition. In conventional Dixon FSE imaging, the refocus andacquisition pulses are carried out repeatedly as indicated by referencenumeral 82 in FIG. 2, wherein successive refocusing pulses 84 areapplied to acquire successive MRI echo signal data, with different phaseencoding for each successive repetition. The number of repetitions willgenerally depend upon the desired image resolution along the phaseencode direction.

[0039] In the conventional FSE pulse sequence, the interecho spacing iskept to a minimum that allows the playout of all the gradient and RFpulses without conflicting pulse overlap. Increased interecho spacinggenerally results in increased total scan time for a given number ofslices, or reduced number of slices within a given total scan time. Inaddition, the increased intercho spacing also leads to a detrimentaleffect of increased image blurring and increased sensitivity to flow andmotion. For the FSE implementation of the Dixon technique illustrated inFIG. 2, the inter-echo spacing must be twice the time period 86 betweenthe 90° excitation pulse 66 and the 180° refocus pulse 70 in order tosatisfy the Carr-Purcell-Meiboom-Gill (CPMG) condition. Moreover, inorder to generate a phase shift between water and fat magnetizations, atime shift 80 is needed, as mentioned above. It has been found, forexample, that in a 1.5 Tesla MRI system, a time 80 required to achieve aphase angle difference of 180° is approximately 2.3 ms. Because theinter-echo spacing must be twice this value, however, a 4.6 ms increasein the echo spacing is required to achieve the desired 180° phase shiftbetween water and fat signals at the 1.5 Tesla field strength. Theseconsiderations result in substantial lengthening of the interechospacing and therefore total data acquisition period. Considering, forexample, a typical 8 ms total readout time when using 256 readout pointsat 16 kHz receiver bandwidth, such an increase amounts to a significantlengthening of the sequence time and in parts increased image blurringdue to the signal T2-modulation within the echo train. More importantly,the increase in echo spacing leads to a substantial protocol-dependentreduction, sometimes as much as 30%-40%, in the slice coverage for agiven imaging time. Again, considering only the issue of the timelengthening, the echo train length is increased by twice the requiredtime offset, multiplied by the number of echoes per excitation for whichdata is desired, typically 12-33.

[0040]FIG. 3 illustrates an improved pulse sequence description forobtaining high-quality Dixon FSE image data, while avoiding theaforementioned added time delay and reducing the image degradationaffects. While reference is made herein to an FSE pulse sequence, itshould be noted that aspects of the present technique are equally wellapplicable to other types of imaging, such as spin echo and gradientecho imaging. Moreover, several alternative examples of the pulsesequence description are available for implementing the technique, asdescribed more fully below.

[0041] Referring to FIG. 3, the pulse sequence description, which may bereferred to as a gradient-induced echo shift sequence 90 is againrepresented by pulses applied along or occurring along rf axis 60,readout gradient axis 62, and data acquisition axis 64. A first rf axisexcitation pulse 66 is applied along the rf axis 60, followed by adephaser gradient 68 applied along the gradient axis 62. A 180° refocuspulse 70 is then applied along the rf axis 60. In the sequence 90 ofFIG. 3, then, a small gradient pulse 92 with minimum duration is addedprior to the readout gradient pulse 72. The net effect of the addedgradient pulse 92 is to induce a spatially linear spin phase shift alongthe readout direction, or equivalently a constant time shift in echoposition. The time shift may be indicated graphically as indicated atreference numeral 94 in FIG. 3, effectively inducing a desired phaseshift between the constituents spin species (e.g. water and fat).

[0042] The area of the shifting gradient pulse 92 is set equal to theproduct of the amplitude of this readout gradient and the desired timeshift in echo position. To preserve the Carr-Purcell-Meiboom-Gill (CPMG)condition, another gradient pulse, which may be referred to as acompensating pulse 96 having the same area but opposite polarity isadded to the sequence after the readout gradient 72. As will beappreciated by those skilled in the art, the shifting gradient pulse 92and compensating pulse 96 may be applied during times when slice crushergradients and phase encode gradients are applied along other logicalaxes. Thus, no timing change in rf, readout gradient, or dataacquisition window locations is necessary. It should also be noted thatsimilar gradient-induced echo shifting gradient pulses 92 are appliedduring subsequent acquisition repetitions, along with compensatinggradient pulses 96. As will be appreciated by those skilled in the art,each refocus fills one line of k-space, or unprocessed data which willbe processed and transformed for image reconstruction. The dataacquisition, of course, proceeds through acquisition repetitions withvarying pre-selected phase shifts (for example, 0, 90 and 180 degrees)that are interleaved with one another. The number of lines of k-space sofilled will depend upon the desired image resolution, with 256 linestypically being filled during an examination.

[0043] In a present implementation, the spatially linear phase shiftalong the readout direction induced by the added gradients iseffectively restored by re-centering the acquired data before imagereconstruction. Consequently, neither increase in the echo spacing norloss in slice coverage is anticipated. It should be further noted thatin the implementation of FIG. 3, an arbitrary echo shift can be effectedsimply by changing the areas of the gradient pulses 92 and 96 withoutchanging the timing of any other pulses in the sequence. Indeed, with afixed pulse width for pulses 92 and 96, echo shift is directlyproportional to, and conveniently controlled by the amplitude of thegradient pulses themselves.

[0044]FIG. 4 represents a variation of the pulse sequence illustrated inFIG. 3. As will be noted in FIG. 4, the polarity of the echo shift pulsegradients 92 and of the compensating gradient 96 may be inverted, whiletheir mutually opposite polarity is maintained. Other aspects of thesequence of FIG. 4 are identical of FIG. 3.

[0045] Although the proposed pulse sequence is compatible with variousmulti-point Dixon acquisition techniques, a three-point asymmetricsampling scheme with 0, 90° and 180° phase shifts was implemented in apresent exemplary embodiment. Apart from increasing timing flexibility,the asymmetric sampling has been shown to offer increased processingreliability. To restore the spatially linear phase shift induced by theadded gradients, echoes with non-zero phase shifts are first re-centeredby an amount equal to the echo shifts induced by the shifting gradientpulses 92 prior to image reconstruction. As a result, the raw data forthese echoes are symmetric with a missing portion that is proportionalto the phase shifts between the desired water and fat magnetizations inthe present exemplary embodiment. In the present exemplary embodiment, ahomodyne reconstruction algorithm was used, but with image phaseinformation being preserved by first performing a Fourier transform inaccordance with the relationship: $\begin{matrix}\begin{matrix}{{I\left( {x,y} \right)} = {^{- 1}\left\{ {{L\left( {k_{x},k_{y}} \right)} + {2{u\left( k_{x} \right)}{H\left( {k_{x},k_{y}} \right)}}} \right\}}} \\{= {{{m\left( {x,y} \right)}\quad ^{{\varphi}{({x,y})}}} + {{h\left( {x,y} \right)}^{{\varphi}{({x,y})}}*\frac{1}{{\pi}\quad x}}}}\end{matrix} & {{Eq}.\quad 1}\end{matrix}$

[0046] where, L(k_(x), k_(y)) and H(k_(x), k_(y)) represent the centralsymmetric and outer asymmetric portions of the acquired data,respectively. The term u(k_(x)) is a unit step function that is used toeffectively double the weight of the asymmetric portion of the data inthe Fourier transform. The term Φ(x,y) is the spatially varying phaseerror that arises from factors such as the gradient timing and RFimperfections. The terms m(x,y) and h(x,y) are the desired and thehigh-resolution components of the image, respectively. The relationshipis valid so long as Φ(x,y) can depict adequately the phase of the imagesreconstructed either using the full or only the central portions of thek-space data. It can be noted that in addition to the desired imagem(x,y)e^(iΦ(x,y)), there is generally a blurring term that is equal tothe convolution of h(x,y)e^(iΦ(x,y)) with a kernel 1/iπx. Thecontribution of the blurring component is related to the amount of themissing portion of the acquired data. When the phase shift is 0, theacquired echo becomes symmetric and the above relationship is reduced toa conventional Fourier transform.

[0047] Because the phase error term e^(iΦ(x,y)) is usually slow-varyingspatially, it can be adequately determined from low-resolution imageswithout compromising the final image resolution. It has beendemonstrated that faster and more reliable determination of the phaseerror terms can be achieved in Dixon processing with low-resolutionimages because they have reduced matrix size and increasedsignal-to-noise ratio. In a present implementation, three sets oflow-resolution images (corresponding to the acquisitions with 0, 90°,and 180° phase shifts) were reconstructed from the central symmetricportion (either 128×128 or 64×64) of the acquired multipoint Dixon data.A region-growing algorithm without direct phase unwrapping, was employedto determine the phase error term e^(iΦ(x,y)) from these threelow-resolution images corresponding to the different echo shifts.

[0048] Assuming e^(iΦ(x,y)) does not vary significantly over the scaleof 1/iπx, it can then be used to demodulate the image from the abovecomputations, in accordance with the relationship: $\begin{matrix}\begin{matrix}{{I_{s}\left( {x,y} \right)} = {{I\left( {x,y} \right)}^{- {{\varphi}{({x,y})}}}}} \\{\approx {{m\left( {x,y} \right)} + {{h\left( {x,y} \right)}*\frac{1}{{\pi}\quad x}}}}\end{matrix} & {{Eq}.\quad 2}\end{matrix}$

[0049] Note that in the regular homodyne reconstruction, both m(x,y) andh(x,y) are assumed to be real, and the blurring component (the secondterm in Eq. 2 above) can be discarded simply by taking the real part ofthe sum. In Dixon imaging, it is recognized that this can also beperformed when the phase angle difference is either 0 or 180° becausethe water and fat magnetizations are then along the same axis. For phaseangles other than 0 or 1800, the object is generally not real and thereis usually an intermix of the blurring component and the desired imagein the real and imaginary channels. Under such circumstances, theblurring component can, in principle, be estimated using knownapproaches, or by iterative methods. In a present implementation,however, because only 36 out of a typical 256 data points were notacquired at 16 kHz receiver bandwidth, simple zero-filling of the dataand direct Fourier transform were used to obtain the image correspondingthe 90° phase shift.

[0050]FIG. 5 represents a further alternative pulse sequence 98 inaccordance with aspects of the present technique which can be used toobtain the desired echo shift. As shown in FIG. 5, the pulses appliedalong the rf axis are essentially similar to those of FIGS. 3 and 4.However, rather than applying the echo shifting gradient pulse 92 priorto the readout pulse 72, an echo shifting gradient pulse 100 is added tothe dephaser pulse 68 applied along the readout axis. To compensate forthe shift induced by the additional gradient pulse 100, a compensatingpulse 102 is applied after the readout pulse 72. Again, the area of thecompensating pulse 102 is essentially the same as that of the addedpulse portion 100. A similar shift in echo timing is thereby induced asindicated at reference numeral 94. As a variation on the pulse sequence98 of FIG. 5, the polarity of the shifting of pulse 100 and of thecompensating pulse 102 may, as in the previous example, be inverted,thereby altering the manner in which the signal time shift occurs, whilemaintianing the desired shift.

[0051] It should be noted that, due to the assymetrical positioning ofthe time-shifted MR echo resulting from the gradient pulses 92, not onlywill one side of the acquired signal be shorter than would be the casewithout such shifting, the opposite side of the signal will be longer orextended beyond the non-shifted length. Where desired, the additionalinformation provided by the extended portion of the signal may be usedto reconstruct images of higher resolution.

[0052] It should be noted that aspects of the foregoing techniques maybe applied somewhat differently to obtain desired echo shifting in othercontexts. For example, other pulse sequences may benefit from theshifting of the MR echo, such as spin echo and gradient echo sequences.In such cases, the negative compensating gradients, while essential formaintaining the CPMG-condition for FSE-based technique, can be left outto further reduce the minimim time required to play out the pulsesequence. In the same spirit, the gradient lobes added for shifting theecho position can possibly be combined with any other gradients on thereadout axis if doing so does not change substantively the spin phaseaccumulation.

[0053] While the invention may be susceptible to various modificationsand alternative forms, specific embodiments have been shown by way ofexample in the drawings and have been described in detail herein.However, it should be understood that the invention is not intended tobe limited to the particular forms disclosed. Rather, the invention isto cover all modifications, equivalents, and alternatives falling withinthe spirit and scope of the invention as defined by the followingappended claims. By way of example, as noted above, the presenttechnique has been described in conjunction with a fast spin echo pulsesequence description. As noted above, however, aspects of the techniquemay be applied beneficially with other pulse sequence descriptions, suchas spin echo imaging sequences. Similarly, while the particular 3-pointasymmetric scheme and the particular image reconstruction techniquedescribed above has been employed in exemplary embodiments, otherschemes and reconstructions techniques may also be used with theimproved pulse sequences.

What is claimed is:
 1. A method for acquiring magnetic resonance imagedata comprising: in the presence of a primary and gradient magneticfield system, generating an echo shifting gradient pulse on a readoutaxis; generating a readout gradient pulse on the readout axis; anddetecting magnetic resonance echo signals resulting from the readoutgradient
 2. The method of claim 1, further comprising generating acompensating gradient pulse on the readout axis.
 3. The method of claim2, wherein the compensating gradient pulse has a polarity opposite apolarity of the echo shifting gradient pulse.
 4. The method of claim 1,wherein the echo shifting gradient pulse produces a spatially linearspin phase shift along a readout direction.
 5. The method of claim 1,wherein the echo shifting gradient pulse produces a constant time shiftin a position of the echo signals.
 6. The method of claim 5, wherein thearea of the echo shifting gradient pulse is substantially equal to theproduct of an amplitude of the readout gradient and a desired time shiftin echo signal position.
 7. The method of claim 1, wherein the echoshifting gradient pulse is generated during a slice crusher gradientpulse.
 8. The method of claim 1, wherein the echo shifting gradientpulse is generated during a phase encode gradient pulse.
 9. The methodof claim 1, wherein the steps are performed in a Dixon fast spin echopulse sequence.
 10. The method of claim 1, comprising repeating thesteps for a series of k-space acquisition lines.
 11. The method of claim1, further comprising reconstructing an image based upon the detectedecho signal data in accordance with a phase-sensitive partial Fourierreconstruction algorithm.
 12. A method for acquiring magnetic resonanceimage data comprising: in the presence of a primary and gradientmagnetic field system, generating an echo shifting gradient pulse on areadout axis; generating a readout gradient pulse on the readout axis;detecting magnetic resonance echo signals resulting from the readoutgradient; and generating a compensating gradient pulse on the readoutaxis.
 13. The method of claim 12, wherein the echo shifting gradientpulse produces a spatially linear spin phase shift along a readoutdirection.
 14. The method of claim 12, wherein the echo shiftinggradient pulse produces a constant time shift in a position of the echosignals.
 15. The method of claim 14, wherein the area of the echoshifting gradient pulse is substantially equal to the product of anamplitude of the readout gradient and a desired time shift in echosignal position.
 16. The method of claim 12, wherein the compensatinggradient pulse has a polarity opposite a polarity of the echo shiftinggradient pulse.
 17. The method of claim 12, wherein the echo shiftinggradient pulse is generated during a slice crusher gradient pulse. 18.The method of claim 12, wherein the echo shifting gradient pulse isgenerated during a phase encode gradient pulse.
 19. The method of claim12, wherein the steps are performed in a Dixon fast spin echo pulsesequence.
 20. The method of claim 12, comprising repeating the steps fora series of k-space acquisition lines.
 21. The method of claim 12,further comprising reconstructing an image based upon the detected echosignal data in accordance with a phase sensitive partial Fourierreconstruction algorithm.
 22. The method of claim 21, comprisingre-centering echo signal data prior to image reconstruction.
 23. Themethod of claim 21, comprising reconstructing a water-only image. 24.The method of claim 23, comprising reconstructing a fat-only image. 25.A method for acquiring magnetic resonance image data comprising: in thepresence of a primary and gradient magnetic field system, applying aDixon pulse sequence; generating an echo shifting gradient pulse on areadout axis; generating a readout gradient pulse on the readout axis;detecting magnetic resonance echo signals resulting from the readoutgradient; and generating a compensating gradient pulse on the readoutaxis.
 26. The method of claim 25, wherein the pulse sequence is a fastspin echo pulse sequence.
 27. The method of claim 25, wherein the echoshifting gradient pulse produces a spatially linear spin phase shiftalong a readout direction.
 28. The method of claim 25, wherein the echoshifting gradient pulse produces a constant time shift in a position ofthe echo signals.
 29. The method of claim 28, wherein the area of theecho shifting gradient pulse is substantially equal to the product of anamplitude of the readout gradient and a desired time shift in echosignal position.
 30. The method of claim 25, wherein the compensatinggradient pulse has a polarity opposite a polarity of the echo shiftinggradient pulse.
 31. The method of claim 25, wherein the echo shiftinggradient pulse is generated during a slice crusher gradient pulse. 32.The method of claim 25, wherein the echo shifting gradient pulse isgenerated during a phase encode gradient pulse.
 33. The method of claim25, comprising repeating the steps for a series of k-space acquisitionlines.
 34. The method of claim 25, further comprising reconstructing animage based upon the detected echo signals in accordance with a phasesensitive partial Fourier reconstruction algorithm.
 35. The method ofclaim 34, comprising re-centering echo signal data prior to imagereconstruction.
 36. The method of claim 34, comprising reconstructing awater-only image.
 37. The method of claim 34, comprising reconstructinga fat-only image.
 38. A method for acquiring magnetic resonance imagedata comprising: in the presence of a primary and gradient magneticfield system, applying a Dixon fast spin echo pulse sequence; generatingan echo shifting gradient pulse on a readout axis; generating a readoutgradient pulse on the readout axis; detecting magnetic resonance echosignals resulting from the readout gradient; and generating acompensating gradient pulse on the readout axis.
 39. The method of claim38, wherein the echo shifting gradient pulse produces a spatially linearspin phase shift along a readout direction.
 40. The method of claim 38,wherein the echo shifting gradient pulse produces a constant time shiftin a position of the echo signals.
 41. The method of claim 40, whereinthe area of the echo shifting gradient pulse is substantially equal tothe product of an amplitude of the readout gradient and a desired timeshift in echo signal position.
 42. The method of claim 38, wherein thecompensating gradient pulse has a polarity opposite a polarity of theecho shifting gradient pulse.
 43. The method of claim 38, wherein theecho shifting gradient pulse is generated during a slice crushergradient pulse.
 44. The method of claim 38, wherein the echo shiftinggradient pulse is generated during a phase encode gradient pulse. 45.The method of claim 38, comprising repeating the steps for a series ofk-space acquisition lines.
 46. The method of claim 38, furthercomprising reconstructing an image based upon the detected echo signalsin accordance with a phase sensitive partial Fourier reconstructionalgorithm.
 47. The method of claim 46, comprising re-centering echosignal data prior to image reconstruction.
 48. The method of claim 46,comprising reconstructing a water-only image.
 49. The method of claim48, comprising reconstructing a fat-only image.
 50. A method foracquiring magnetic resonance image data comprising: in the presence of aprimary and gradient magnetic field system, applying a Dixon fast spinecho pulse sequence to acquire a plurality of k-space lines of data; foreach k-space line of data: generating an echo shifting gradient pulse ofa first polarity and of a desired area on a readout axis; generating areadout gradient pulse on the readout axis; detecting magnetic resonanceecho signals resulting from the readout gradient; and generating acompensating gradient pulse of a second polarity opposite to the firstpolarity and of the desired area on the readout axis.
 51. The method ofclaim 50, wherein each echo shifting gradient pulse produces a spatiallylinear spin phase shift along a readout direction.
 52. The method ofclaim 50, wherein each echo shifting gradient pulse produces a constanttime shift in a position of the echo signals.
 53. The method of claim50, wherein the desired area of the echo shifting gradient pulse issubstantially equal to the product of an amplitude of the readoutgradient and a desired time shift in the echo signal position.
 54. Themethod of claim 50, wherein each echo shifting gradient pulse isgenerated during a slice crusher gradient pulse.
 55. The method of claim50, wherein each echo shifting gradient pulse is generated during aphase encode gradient pulse.
 56. The method of claim 50, furthercomprising reconstructing an image based upon the detected echo signalsin accordance with a phase sensitive partial Fourier reconstructionalgorithm.
 57. The method of claim 56, comprising re-centering thedetected signal data prior to image reconstruction.
 58. The method ofclaim 56, comprising reconstructing a water-only image.
 59. The methodof claim 58, comprising reconstructing a fat-only image.
 60. A methodfor acquiring magnetic resonance image data comprising: in the presenceof a primary and gradient magnetic field system, applying a Dixon fastspin echo pulse sequence to acquire a plurality of k-space lines ofdata; for each k-space line of data: generating an echo shiftinggradient pulse of a first polarity and of a desired area on a readoutaxis; generating a readout gradient pulse on the readout axis; detectingmagnetic resonance echo signals resulting from the readout gradient;generating a compensating gradient pulse of a second polarity oppositeto the first polarity and of the desired area on the readout axis; andreconstructing an image based upon the detected echo signals inaccordance with a phase sensitive partial Fourier reconstructionalgorithm.
 61. The method of claim 60, wherein each echo shiftinggradient pulse produces a spatially linear spin phase shift along areadout direction.
 62. The method of claim 60, wherein each echoshifting gradient pulse produces a constant time shift in a position ofthe echo signals.
 63. The method of claim 60, wherein the *desired areaof the echo shifting gradient pulse is substantially equal to theproduct of an amplitude of the readout gradient and a desired time shiftin the echo signal position.
 64. The method of claim 60, wherein eachecho shifting gradient pulse is generated during a slice crushergradient pulse.
 65. The method of claim 60, wherein each echo shiftinggradient pulse is generated during a phase encode gradient pulse. 66.The method of claim 60, comprising re-centering the detected signal dataprior to image reconstruction.
 67. The method of claim 60, comprisingreconstructing a water-only image.
 68. The method of claim 60,comprising reconstructing a fat-only image.
 69. An image generated bythe method of claim
 60. 70. A magnetic resonance imaging systemcomprising: means for applying a Dixon fast spin echo pulse sequence toacquire a plurality of k-space lines of data; means for generating anecho shifting gradient pulse of a first polarity and of a desired areaon a readout axis; means for generating a readout gradient pulse on thereadout axis; means for detecting magnetic resonance echo signalsresulting from the readout gradient; means for generating a compensatinggradient pulse of a second polarity opposite to the first polarity andof the desired area on the readout axis; and means for reconstructing animage based upon the detected echo signals.